The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to flow compensation of steady-state free precession pulse sequences.
Any nucleus which possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the magnetogyric constant .gamma. of the nucleus). Nuclei which exhibit this phenomena are referred to herein as "spins".
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B.sub.0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net magnetic moment M.sub.z is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or traverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B.sub.1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, M.sub.z may be rotated, or "tipped", into the x-y plane to produce a net transverse magnetic moment M.sub.t, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The degree to which the net magnetic moment M.sub.z is tipped, and hence the magnitude of the net transverse magnetic moment M.sub.t depends primarily on the length of time and the magnitude of the applied excitation field B.sub.1.
The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation signal B.sub.1 is terminated. In simple systems the excited spin induce an oscillating sine wave signal in a receiving coil. The frequency of this signal is the Larmor frequency, and its initial amplitude, A.sub.0, is determined by the magnitude of the transverse magnetic moment M.sub.t. The amplitude, A, of the emission signal decays in an exponential fashion with time, t: EQU A=A.sub.0 e.sup.-t/T.spsp.*.sbsp.2
The decay constant 1/T.sup.*.sub.2 depends on the homogeneity of the magnetic field and on T.sub.2, which is referred to as the "spin-spin relaxation" constant, or the "transverse relaxation" constant. The T.sub.2 constant is inversely proportional to the exponential rate at which the aligned precession of the spins would dephase after removal of the excitation signal B.sub.1 in a perfectly homogeneous field.
Another important factor which contributes to the amplitude A of the NMR signal is referred to as the spin-lattice relaxation process which is characterized by the time constant T.sub.1. It describes the recovery of the net magnetic moment M to its equilibrium value along the axis of magnetic polarization (z). The T.sub.1 time constant is longer than T.sub.2, much longer in most substances of medical interest.
The NMR measurements of particular relevance to the present invention are called "pulse NMR measurements". Such NMR measurements are divided into a period of RF excitation and a period of signal emission. Such measurements are performed in a cyclic manner in which the NMR measurement is repeated many times to accumulate different data during each cycle or to make the same measurement at different locations in the subject. A wide variety of preparative excitation techniques are known which involve the application of one or more RF excitation pulses (B.sub.1) of varying magnitude, duration, and direction. Such excitation pulses may have a narrow frequency spectrum (selective excitation pulse), or they may have a broad frequency spectrum (nonselective excitation pulse) which produces transverse magnetization M.sub.t over a range of resonant frequencies. The prior art is replete with excitation techniques that are designed to take advantage of particular NMR phenomena and which overcome particular problems in the NMR measurement process.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (G.sub.x, G.sub.y, and G.sub.z) which have the same direction as the polarizing field B.sub.0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified. NMR data for constructing images can be collected using one of many available techniques. Typically, such techniques comprise a pulse sequence made up of a plurality of sequentially implemented views. Each view may include one or more NMR experiments, each of which comprises at least an RF excitation pulse and a magnetic field gradient pulse to encode spatial information into the resulting NMR signal.
Most NMR scans currently used to produce medical images require many minutes to acquire the necessary data. The reduction of this scan time is an important consideration, since reduced scan time increases patient throughput, improves patient comfort, and improves image quality by reducing motion artifacts. There is a class of pulse sequences which have a very short repetition time (TR) and result in complete scans which can be conducted in seconds rather than minutes. Whereas the more conventional pulse sequences have repetition times TR which are much greater than the spin-spin relaxation constant T.sub.2 so that the transverse magnetization has time to relax between the phase coherent excitation pulses in successive sequences, the fast pulse sequences have a repetition time TR which is less than T.sub.2 and which drives the transverse magnetization into a steady-state of equilibrium. Such techniques are referred to as steady-state free precession (SSFP) techniques and they are characterized by a cyclic pattern of transverse magnetization in which the resulting NMR signal refocuses at each RF excitation pulse to produce an echo signal. This echo signal includes a first part S+ that is produced after each RF excitation pulse and a second part S- which forms just prior to the RF excitation pulse.
There are two well known SSFP pulse sequences used to produce images. The first is called gradient refocused acquired steady-state (GRASS) and it utilizes a readout gradient G.sub.x to shift the peak in the S+ signal that is produced after each RF excitation pulse toward the center of the pulse sequence. This pulse sequence is shown in FIG. 1 where the NMR signal is an S+ gradient echo that is induced by the readout gradient G.sub.x. In two-dimensional imaging, a slice selection gradient pulse is produced by the gradient G.sub.z and is immediately refocused in the well-known manner. A phase encoding gradient pulse G.sub.y is produced shortly thereafter to position encode the acquired NMR data, and to preserve the steady-state equilibrium. The effects of the phase encoding gradient pulse are nullified by a corresponding G.sub.y rewinder gradient pulse after the NMR signal has been acquired and before the next pulse sequence begins as described in U.S. Pat. No. 4,665,365.
The second well known SSFP pulse sequence is called contrast enhanced fast imaging (SSFP-ECHO) and it utilizes the S- signal that is produced just prior to each RF excitation pulse. This pulse sequence is shown in FIG. 2 where the acquired NMR signal is an S- echo signal caused by the gradient refocusing of the transverse magnetization which would otherwise refocus at the next RF excitation pulse. The readout gradient G.sub.x is substantially different in this pulse sequence and includes a positive pulse prior to the actual readout pulse and a negative pulse after the readout pulse. The former pulse dephases the FID signal (S+) which might otherwise be produced during the data acquisition window, and the latter pulse causes the transverse magnetization to rephase during the next pulse sequence to produce the echo signal S- during the acquisition window. For a more detailed discussion of the SSFP-ECHO or S- pulse sequence, reference is made to an article by R. C. Hawkes and S. Patz entitled "Rapid Fourier Imaging Using Steady-State Free Precision", published in Magnetic Resonance in Medicine 4, pp. 9-23 (1987).
In addition to being a very short pulse sequence which enables a complete scan to be carried out in a few seconds, the S- sequence has an attribute which makes it more useful than the GRASS pulse sequence in many medical applications. More specifically, the S+ signal acquired with the GRASS pulse sequence has an amplitude which is approximately a function of the ration T.sub.2 /T.sub.1, while the S- signal acquired by the S- pulse sequence has additional T.sub.2 dependence. As a result, S- is a T.sub.2 weighted pulse sequence which provides better contrast between tissues of differing T.sub.2. Since T.sub.2 is a good indicator of diseased tissues, the S- pulse sequence provides better contrast between normal tissues and diseased tissues in the reconstructed image.
Unfortunately, the S- pulse sequence also is more susceptible to distortions, or artifacts, which are produced in the reconstructed image as a result of moving spins. For example, blood flow in the direction of the readout gradient is incorrectly position encoded, and the signals produced by the flowing spins are thus positioned incorrectly in the reconstructed image. This increased sensitivity to motion and flow artifacts is due to the fact that the S- signal is produced by transverse magnetization which is created by RF excitation pulses generated during previous pulse sequences. Over the extended time (&gt;2TR) between excitation and readout, the gradient fields produce large dipolar moments which sensitize the S- signal to flow and motion.
In my prior invention which is disclosed in U.S. Pat. No. 4,973,906 entitled "Flow Compensated NMR Fast Pulse Sequence," I disclose a S- pulse sequence which is flow compensated. Referring to FIG. 3 in this application, the G.sub.x readout gradient pulse in this prior art invention is comprised of a central positive gradient pulse 10, a leading negative gradient pulse 11, and a trailing negative gradient pulse 12. The readout pulse 10 refocuses the S- NMR signal and it frequency encodes the acquired data in the well known manner. The trailing negative gradient pulse 12 dephases the spins and sets up the production of the S- NMR signal during the next pulse sequence. A teaching of this prior invention is that the leading negative gradient pulse 11 is chosen so as to flow compensate the readout gradient and thereby reduce the production of flow artifacts. In practice this compensation cannot be accomplished completely and a compromise value is used.